The optimal detection of ionizing radiation in two dimensions is the central problem in computed tomography, digital radiography, nuclear medicine imaging and related disciplines. Many different types of detectors (e.g., non-electronic, analog electronic and digital electronic detectors) have been used with varying degrees of success in these fields. In general, many compromises have been made in the various imaging and non-imaging parameters of detectors in developing opertional systems.
More recently, there has been developed a different type of detector known as the kinestatic charge detector (KCD). In a KCD system, there is provided an x-ray detection volume and a signal collection volume formed in a closed chamber. In the etection volume, there is generally disposed some type of medium which will interact with x-ray radiation to produce secondary energy. The medium is generally enclosed within a defined space and the collection volume is preferably a multi-element detector of secondary energy located at one boundary of the detection volume. An applied electric field across the detection volume imparts a constant drift velocity to secondary energy particles or charges driving the charges of one sign towards the signal collection volume. Charges of the other sign will drift in a direction away from the collection volume and will not contribute to any output signal.
In the operation of the system, an x-ray beam scans a patient and the x-ray radiation passing through the patient is directed into the detection volume. The x-ray radiation collides with particles in the medium of the detection volume creating a secondary energy. The electric field across the detection volume is produced between a first electrode at one side of the detection volume and the plane of the collection volume (collection electrodes) and the direction of the field is substantially perpendicular to the path of the radiation admitted into the detection volume. The electric field causes charge carriers between the first electrode and the collection electrode to drift toward the collection electrode at a substantially constant drift velocity. The chamber itself, including the detection and collection volumes, is mechanically coupled to apparatus which moves the chamber in a direction opposite to the direction of drift of the charges at a constant velocity of a magnitude substantially equal to the magnitude of the drift velocity of the charges. The currents flowing in the plural collection electrodes resulting from charges produced on the collection electrodes by the charge carriers is sensed. The spatial distribution in two dimensions of the radiation admitted into the chamber is determined in response to the amplitude with respect to time of the sensed current flowing in the respective plural collection electrodes.
Since the motion of the chamber is in a direction opposite to the drift of the gas ions created in the medium in the detection volume, the x-ray radiation passing through each small area of the patient in the x-ray beam is integrated over the time that it takes for the ions in the detection volume to drift through the space of the volume. In essence, the motion of the detector combined with the motion of the particles combine to make the x-ray radiation appear to be stationary with respect to the drifting particles. Within the detection volume, a grid is required to separate the space between the first electrode and the collector volume into a drift region and a collection region. The grid shields the collector electrodes from the induced current caused by the charges drifting in the drift region. Since the grid and collector electrodes are at different electrical potentials, the electrodes will be sensitive to microphonic noise caused by relative motion between the grid and the collector electrodes. Such microphonic noise may be caused by motion of the chamber or by other external vibrations induced into the support structure for the chamber. The microphonic noise will result in inaccurate detection of the charged particles and in reproduction of an inaccurate presentation of the actual image of the patient.
The production of microphonic noise by relative motion between a grid and electrodes in an ionization chamber is recognized in U.S. Pat. No. 4,047,040 issued Sept. 6, 1977 and assigned to General Electric Company, although that patent discloses a system in which noise is generated by motion of the anode and cathode electrodes rather than motion of the grid structure. In that patient, an ionzation chamber for a computerized tomography system is illustrated. The microphonic problem is resolved by attaching the grid directly to the anode through an insulating material, the insulating material being deposited on the anode structure. Because the grid need only be maintained at a 30 volt differential with respect to the anode, such a solution is satisfactory. However, as pointed out in that patent, even with an insulative layer of about 0.1 millimeter, the grid structure reduces the quantity of electrons reaching the anode by nearly 50 percent. Such a degree of attenuation is unsatisfactory in a KCD system. Furthermore, in a KCD system, the voltages required are orders of magnitude greater than the CT voltages and would require a substantial increase in the required electrical resistivity of the insulator (typically 10.sup.14 ohms or greater) to reduce electrical leakage from the grid to the collector to a satisfactory level.
It is an object of the present invention to provide an improved grid structure for an ionization chamber.
It is a still further object of the present invention to provide an improved grid structure for use in a KCD system.